In the field of nuclear medical imaging technology, a subject, e.g. an oncology patient or an animal used in an experiment, can be scanned by detecting radiation emanating from the subject. For example, in a so-called PET scan, a short-lived radioisotope, which decays by emitting a positron, is injected usually into the blood circulation of a living subject. After the metabolically active molecule becomes concentrated in tissues of interest, the research subject or patient is placed in the imaging scanner. The most commonly-used metabolically active molecule for this purpose is 18F-fluorodeoxyglucose (FDG), a sugar, which has a half life of 110 minutes.
As the radioisotope undergoes positron emission decay, it emits a positron, the antimatter counterpart of an electron. After traveling up to a few millimeters, the positron encounters and annihilates with an electron, producing a pair of gamma photons moving in almost opposite directions. These are detected when they reach one of a plurality of scintillation crystals in the scanning device, creating a burst of light detected by an array of photosensors.
Radiation emanating from the subject can be detected in, for example, radiation detector ring assembly 106 illustrated in FIG. 1. At a more granular level, specific radiation events can be detected at detector block 112 comprising an array of radiation sensors, such as plurality of scintillators and associated photosensors 102n, such as photomultiplier tubes (PMTs), avalanche photodiodes (APDs), or silicon photomultipliers (SiPMs). In the case of a PET scan, scintillators can be arranged in a ring 106.
Generally a plurality of sensors, e.g., photosensors 102, can be arranged in a matrix and assigned to detect the light of a scintillator as shown in detector block 112 in FIG. 1. Detector block 112 can be associated with a single scintillation crystal 106 or can be, as shown, a matrix of scintillator crystals that is coupled to the photosensors 1021 . . . 102n usually via a light guide. A plurality of detector blocks 112 can be axially arranged adjacent to one another, in a slot, in a line relative to the center of ring 106. To be able to increase the resolution of the system without the high costs of 1:1 coupling, the number of photosensors 102 per block is generally significantly lower than the number of scintillation crystals 106. For example, a detector block may have a plurality of radiation sensors, such as photosensors 102 with, for example, 4, 9 or 16 photosensors 102 arranged in a 2×2, 3×3, or 4×4 matrix behind an array of scintillation crystals 106. Other arrangements with more or fewer photosensors 102 are possible. Thus, scintillation event localization can be determined or interpolated by such a detector block by processing the associated photosensor signals. This can be done by analog filtering, integration, and multiplication of weighted combinations of the photosensor signals or by using digital algorithms that use discrete time sample points of signals obtained directly from the photosensors 102. The PET technique depends on scintillation event detection of the pair of gamma photons.
FIG. 1 illustrates a block diagram of the typical architecture of a detector block 112 and associated analog-to-digital-converters 108-108n in a conventional system. Each matrix of photosensors 102 produces a plurality of signals that can be processed to generate an image from a plurality of scintillation events that are detected by a photosensor 102. To determine the location of a detected annihilation, the system needs to accurately measure the timing and energy of both detected photons. Consequently a high amount of data has to be produced by the respective measurement circuits.
For example, as shown on the right side of FIG. 1, each scintillator has an associated matrix of detector blocks, such as photosensors 1021 . . . 102n, which, in this example are PMTs. Each signal of each PMT 1021 . . . 102n is first amplified by, for example, associated preamplifiers/buffers 1041 . . . 104n. The output signal of preamplifier/buffers 1041 . . . 104n can then be converted concurrently into discrete-time digital signals by associated analog-to-digital converters (ADC) 1081 . . . 108n. A sampling clock for each ADC can be provided at terminal 110. In this example, this digital processing architecture uses n independent ADC signals with peripheral circuitry to concurrently sample each of n photosensor signals per block. This can increase the costs of a detector block.
Not all radiation emanating from a subject is detected by scanner 100. Radiation can be emitted outside of the field of view of scanner 100, or radiation can scatter. For example, Compton scatter can occur when a photon collides with an electron, thereby transferring energy to the electron. The collision can cause the photon to deviate from its original path and cause a loss of energy. This collision typically occurs within the subject or in, for example, a scintillation crystal. Due to Compton scattering, events that would otherwise have been detected may be missed. Techniques are known, however, to determine whether a detected gamma photon is a Compton scatter photon and to calculate its original direction to within a certain probability.
The probability that a 511 keV gamma ray be detected is a function of the material composition of the detector block, its size, and its density. For LSO, the probability that the first interaction of the 511 keV gamma ray is a Compton scatter is on the order of 68%, and for short, narrow pixels, the fraction of Compton scatter exiting the pixel can be quite significant.
In addition, the probability of detecting any particular photon from the scintillator depends on the photosensor, for example a SiPM. Each SiPM photosensor pixel consists of a plurality of cells (sometimes also called microcells) which contribute to an overall, summed signal of the photosensor pixel. Each cell is a small avalanche photodiode operating in Geiger mode, above breakdown. When too many photons are captured by a single photosensor pixel, the number of impinging photons can potentially equal or exceed the number of available cells on the photosensor pixel. Because the cells detect a single photon, encountering photons equal to or greater than the number of cells triggers the cells to excess. In this state the photosensor pixel cannot distinguish between one or more of the simultaneously impinging photons, resulting in degradation of signal linearity. In addition, this effect creates additional statistical noise contribution, leading to losses in energy resolution and also in time resolution of the signal.
FIG. 2 illustrates a photosensor pixel 102 comprising a uniform n×n array of photosensor cells; for example, each cell can be 50 μm×50 μm. The photosensor pixel 102 is part of an array of pixels 102-102n, which is coupled, for example, to one or more scintillation crystals 106 to detect light emitted from the crystal(s) due to scintillation events.
FIG. 3 illustrates an example photosensor pixel 300 having uniform cell geometry receiving an asysmmetric amount of light 302. In this example, the signal will saturate the cells in the photosensor pixel much sooner in the bottom right corner than in the remaining portions of the photosensor area, resulting in poor signal linearity. Correction for the asymmetrical amount of light received requires obtaining an optimum trade-off between detection efficiency and signal linearity. Detection efficiency increases with large photosensor cell sizes, because there is usually a minimum gap size between the sensitive parts of the cells, and more area is lost to these gaps as the cells become smaller. On the other hand, the signal linearity increases with higher cell density, because then it is less likely for photons to impinge on the same cells. The optimum trade-off between detection efficiency and signal linearity depends on the specific patterns of light detected. However regardless of the light levels detected, either detection efficiency or signal linearity will be compromised—the exact comprise depends on which trade-off will be more favorable for a particular situation.
FIGS. 4a-4d illustrate a typical detector design for PET imaging. PET detectors are often built as block detectors, where an array of scintillator crystals is coupled to an array of photosensors. In this example, there are 3×3 scintillation crystals coupled to 2×2 sensors, as most clearly seen in FIG. 4a. FIG. 4b illustrates a non-uniform light distribution for emission from the center scintillator pixel, FIG. 4c illustrates emission from an edge pixel, and FIG. 4d illustrates emission from a corner pixel. A result of these light distribution patterns is that the relevant signal range per photosensor pixel for those outer areas is higher than the signal range for the inner areas. In addition, the signal for the corner events will saturate earlier when a uniform cell density is used. While on average the outer areas do not receive more light, the light can be more concentrated on these outer areas for gamma interactions in certain scintillator crystals.